3.1 How do Imaging Scanners and the 3-D Complete Body Scan Work
FIG. 5 shows (a) the evolution of PET instruments in the past several years and (b) includes a comparison to the approach described in this document.
The reduction in radiation dose required to be delivered to the patient, the lower examination cost, the faster scanning time, the better quality image obtained by accumulating more photons in coincidences shown in FIG. 5b (3D-CBS), are provided by the new gantry design and the new approach of the electronics as described in Section 6.2, Section 6.3, and shown in FIG. 16 and FIG. 17. The elimination of the bottleneck on input is described in Section 6.6.7.1.3; the elimination of the bottleneck on output is described in Section 6.6.8.1.4. The elimination of the detector boundaries is described in Section 6.6.7.1.2, and its implementation is shown in FIG. 56; and the elimination of the limitation on the coincidence detection is described in Section 6.6.8, and in Section 6.5.14.
3.2 Solution Needed to Overcome the Efficiency Limitation Imposed by the Architectural Approach of Current Imaging Devices.
3.2.1 Why PET has Not Been Widely Used in the Past 25 Years in Spite of the Excellent, Fast Detectors Available for 10 Years
The advent of PET in the last 25 years has not had a striking impact in hospital practice and has not been widely used because the electronics with the capability of fully exploiting the superiority of the PET technique has never been designed. Currently the best PET detect about 2 photons out of 10,000 (see references [1], and [2]). If used in 2-D mode, they can detect about 2 out of 100,000, while the Single Photon Emission Computed Tomography (SPECT) devices can detect only about 1 out of 200,000 photons (for one head SPECT; and about 1 out of 100,000 for two heads SPECT) emitted by the source.
The aim of this 3D-CBS design is to detect about 1,000 out of 10,000 photons emitted by the source.
Low efficiency in detecting photons without the capability of fully extracting the photon's properties gives poor images that cannot show small tumors, making the device unsuitable for early detection. In addition, it requires high radiation to the patient, which prevents annual examination; and it requires more imaging time, which limits its use to fewer patients per hour, driving the examination cost very high.
The great potential of PET is exploited only if it does not require the use of a lead collimator between the patient and the detector, and if it has an efficient electronics that does not saturate and that fully extracts particle properties using a thorough real-time algorithm.
Conversely, the advances in detector technology have been superb, providing for more than 10 years fast crystals (e.g., LSO with a decay time of the order of 40 ns) and the construction of detectors with small crystals that help to limit to a small area of the detector the dead time of a crystal that received a photon.
3.2.2 Measurements Showing that the Electronics is the Factor Limiting Efficiency in Current PET and Those Under Design
That the electronics is the limiting factor of the efficiency of current PET (besides the plots of PET working in 3-D as described later) is shown by the fact that some PETs currently used in hospitals operate in what is called 2-D mode. 2-D refers to the use of a lead collimator placed in front of the detector. This is used to limit the number of photons hitting the detector (in particular for body scan where Compton scattering is more numerous than in a smaller volume head-scan) because the electronics cannot handle the unregulated rate of photons hitting the detector. The real-time algorithm of current PET cannot thoroughly process all the information necessary to separate a good event from bad events. It is unfortunate that a superior technology such as positron emission is employed in several PETs now in use in hospitals as if it were a SPECT, where the direction of the photons is determined by the holes of a lead collimator. This obviously will prevent many photons not sufficiently aligned with the holes of the collimator from ever reaching the detector.
The saturation of the electronics of current PET, even during levels of low radiation activity; is confirmed in the measurements of the sensitivity reported in the articles of the past 25 years and is graphically represented in a form similar to FIG. 6a. 
The limitation caused by the saturation of the CTI/Siemens electronics [3] (at 10 Mcps), is shown in FIG. 3 of [4]. This is a simulation made by Moses and Huber (see reference [4]) of a PET camera that completely encloses a small animal in a volume formed by 6 planar banks of detector modules. The caption of FIG. 3 of reference [4] says: “The random fraction is small due to the absence of “out of field” activity implicit with complete solid angle coverage, as well as a short coincidence windows. The total scatter event rate is 11% of the total true event rate. A maximum system count rate of 10 Mcps is assumed.” The plots shown in FIG. 3 of [4] are compared with the measurements of the sensitivity of the existing MicroPET with short FOV and thin (10 mm) crystals of the CTI/Siemens [5]. The latter also reveal saturation of the electronics in FIG. 2a of [5].
3.2.3 Efficiency Limitation Imposed by the Architectural Approach of Current Imaging Devices.
After having studied the behavior of the physics experiment in a PET detector we can plot the performance of PETs with different FOV in detecting coincidences vs. the activity of the γ-rays created inside the body, the ones that leaves the body and the ones that hit the detector aperture for different detector FOV.
FIG. 7 shows the plot of the previous graph with the performance of the current PET systems added to it for graphical comparison. The curve at bottom has been calculated from the measurement of the performance of a few of the latest models of whole-body PET systems and/or simulations as reported in recent articles (see the following paragraphs in this section).
In particular one can find on page 115 of reference [6] the description of a PET examination using the model by Siemens ECAT EXACT HR providing an efficiency of only being 0.0193%.
A second reference [2] (on page 1405, FIG. 8) describes a PET examination using GE Advance with the injection of 8.5 mCi 18FDG in a human, yielding a total efficiency of 0.022%.
The efficiency of the most advanced current PET devices is even lower when performance measurements are made using radiotracers such as 15O-water, which generates a higher radiation activity for a shorter time.
For example, the results of the PET brain examination performed with the GE Advance Positron Emission Tomograph on humans using 66 mCi of intravenous injection of the radiotracer 15O-water, yield a total efficiency of 0.0014%. (See reference [2])
CTI/Siemens and General Electric have not proposed increasing the field of view to 120 cm, which would capture most of the radiation delivered to the patient instead of capturing only about 0.022% of the coincidence photons generated, because the current approach that they are using of checking for coincidences on each Line-Of-Response (LOR) would require the number of LOR to increase as the formula ((n×(n−1)/4). Using the current CTI/Siemens and GE approach, the complexity of the electronics would increase enormously, or, alternatively, one would have to drop many photons from being checked. In that case, however, no significant advantage is provided to the patient, because the radiation and the cost have not been lowered.
During these past 25 years, the problem of the electronics has always been considered greater than the benefit which would accrue from the availability of a more efficient PET device. However, a device capable of shortening the examination time would in effect lower the cost per examination, since more patients could be examined each hour. Even more important, lowering the radiation dose to the patient, would enable patients to take the examination more often.
FIG. 7 shows an area where improvements of the current PET devices are necessary, including increasing the Field-Of-View and improving the electronics.
3.3 Deficiencies of Current Medical Imaging Instrumentation.
Although the CT images are of good quality at the expenses of a relatively high x-ray beam (which should be lowered in order to lower the risk to the patient), the PET images are of poor quality because only a few emitted photons from the patient's body are captured by the PET detector. Other deficiencies of the current PET machines are: low coverage of the entire body, false positives, high radiation dose, slow scanning, high examination costs. The increased efficiency of the 3D-CBS in capturing photons, will provide improvements in both: lowering the radiation dosage for CT scan and improve the PET image quality (in addition to also lower PET radiation dosage).
Briefly, following is a list of the main areas of inefficiencies in the current PET which prevent maximum exploitation of positron emission technology.    1. The image quality of current PET is poor because it has:            a. a short FOV, limit by a non efficient electronics that do not offset the cost of the detector if the FOV were increased (see also next section about the false positive and false negatives);        b. no accurate time-stamp assigned to each photon (a) limiting the detection of neighboring photons emitted within a short time interval, (b) causing long dead-time of the electronics and (c) increasing randoms, or photons in time coincidence belonging to two different events, (most PETs do not have any photon time-stamp assignment);        c. analog signal processing on the front-end electronics limiting photon identification because of poor extraction of the characteristics of the incident photon and absence of the capability to improve signal-to-noise (S/N) ratio;        d. detector boundary limitation to 2×2 PMT blocks, no correlation between signals from neighboring detector blocks, no full energy reconstruction of the photons that hit the detector, (most of current PET do not attempt to make any energy reconstruction of the event, but take decisions in accepting or rejecting first a photon and later an event based on the threshold of a single signal).        e. dead-time of the electronics. Dead-time of the electronics is due to any bottleneck (e.g., multiplexing of data from many lines to a single line, saturation on input, processing, saturation on output) present at any stage of the electronics;        f. saturation of the electronics at the input stage due to its inability to detect and process two nearby photons that hit the detector within a short time interval;        g. costly and inefficient coincidence detection circuit (most current PET [20], [18] have a coincidence detection circuit that tests for coincidence all possible combinations of the Lines of Response (LOR) passing through the patient's body). Although current PET have made a compromise in coincidence detection efficiency versus circuit complexity, by using a coarse segmentation of the detector in order to reduce the number of LOR to be tested for coincidence, that approach is however an impediment to increasing the FOV (See more details in Section 14.7.2 of [7] and Section 6.3 of [12]). This approach adds unnecessary complexity to the electronics of the current PET and makes it unreasonably costly to build a circuit with an acceptable efficiency when more detector elements are added to the detector (which is required in extending the FOV);        h. saturation of the electronics at the output stage due to the limiting architecture of the coincidence detection circuit (See Section 5.6.8.1.4);        i. High number of “Randoms” due to the non accurate measurement of the photon arrival time and to the long (about 12 ns) time window used when determining if two photons belong to the same event;        j. Poor measurement of the attenuation of different tissues at different locations in a patient's body. These measurements are necessary for calculating the attenuation correction coefficients for PET scan;            2. The false positives and false negatives shown in images from current PET, are a consequence of all of the above not having: (a) a DSP (see Section 7.2) on each electronic channel, with neighboring signal correlation capabilities, which extracts with zero dead time, the full characteristics of the incident photon and improves the S/N ratio of the each signal before adding it to other signals, (b) good attenuation correction coefficients, (c) a good, efficient, and simple coincidence detection circuit, and (d) a sufficiently long FOV (which prevent capturing most photons as shown on the left side of FIG. 12) that are the impediments in obtaining good quality images;    3. The high radiation dose delivered to the patient is required by the current PET because each examination needs more than 20 million photons in coincidence (or a number that provides a sufficient statistic to build an image). The short FOV and the inefficient electronics allow to accumulate fewer than 2 photons in coincidence every 10,000 emitted. This inefficiency requires to administer necessarily high radiation dosage to the patient in order to keep the examination time within an hour.    4. The slow scanning time is because of the short FOV of the current PET and of the low efficiency of the electronics. The limited efficiency mentioned above of 2 out of 10,000 requires long acquisition time. Examinations longer than one hour are unacceptable because (a) the biological process desired to observe and the radioisotope decay activity would be over, (b) the patient would be uncomfortable, and (c) the cost would be even higher that what it already is;    5. The current high cost of the examination is due to:            the high cost of the huge dose of radioisotope required;        the slow scanning time that allows only six to seven patients per day to be examined; and        the cost of highly paid personnel who must operate the slow machine.        